In CT X-ray imaging of a patient, X-rays are used to image internal structure and features of a region of the patient's body. The imaging is performed by a CT-imaging system, hereinafter referred to as a “CT-imager”, which images internal structure and features of a plurality of contiguous relatively thin planar slices of the body region using X-rays.
The CT-imager generally comprises an X-ray source that provides a planar, fan-shaped X-ray beam and an array of X-ray detectors that are coplanar with the fan beam and face the X-ray source. The X-ray source and array of detectors are mounted in a gantry so that a patient being imaged with the imager, generally lying on an appropriate support couch, can be positioned in a space within the gantry between the X-ray source and the array of detectors. The gantry and couch are moveable relative to each other so that the X-ray source and detector array can be positioned axially at desired locations along the patient's body.
The gantry comprises a stationary structure referred to as a stator and a rotary element, referred to as a rotor, which is mounted to the stator so that the rotor is rotatable about the axial direction. Angular position of the rotor about the axial direction is controllable so that the X-ray source can be positioned at desired angles, referred to as “view angles”, around the patient's body.
To image a slice in a region of a patient's body, the X-ray source is positioned at the axial position of the slice and the X-ray source is rotated around the slice to illuminate the slice with X-rays from a plurality of different view angles. At each view angle, detectors in the array of detectors generate signals that are measures of intensities of X-rays from the source that pass through the slice. The intensity of X-rays measured by a particular detector in the array of detectors is a function of an amount by which X-rays are attenuated by material in the slice along a path length from the X-ray source, through the slice, to the particular detector. The measurement provides information on composition and density of tissue in the slice along the path-length.
For example, if incident X-ray intensity sensed by an “n-th” detector in the array of detectors when the X-ray source is located at a view angle θ is represented by I(n,θ), then I(n,θ)=Ioexp(−∫μ(l)dl). In the expression for I(n,θ), Io is intensity of X-rays with which the X-ray source illuminates the slice, integration over l represents integration over a path through material in the slice along a direction from the X-ray source to the n-th detector and μ(l) is an absorption coefficient for X-rays per unit path-length in the material at position l along the path. (Dependence of the integral on n and θ is not shown explicitly and is determined through dependence of the length and direction of the path-length l on n and θ.) Intensity Io of X-rays from the X-ray source is generally monitored by a “reference detector”, usually located near the X-ray source.
From measured values of Io and I(n,θ) an amount by which X-rays are attenuated along path-length l and a value for the integral ∫μ(l)dl, hereinafter referred to as an “absorption integral”, can be determined. The attenuation measurement provided by the n-th detector at the view angle θ therefore provides a value for the line integral of the absorption coefficient along a particular path length through the slice, which is determined by θ and the known position of the n-th detector relative to the X-ray source.
The set of attenuation measurements for a slice provided by all the detectors in the detector array at a particular view angle θ is referred to as a view. The set of attenuation measurements from all the views of the slice is referred to as a “projection” of the slice. Values for the absorption integral provided by data from the projection of the slice are processed using algorithms known in the art to provide a map of the absorption coefficient μ as a function of position in the slice. Maps of the absorption coefficient for the plurality of contiguous slices in the region of the patient's body are used to display and identify internal organs and features of the region.
In some CT-imagers, to image a region of a patient, the region is scanned by moving the patient stepwise in the z direction to “step” the region through the space inside the CT-imager's gantry between the imager's X-ray source and detector array. Following each step, the X-ray source is rotated through 360 degrees or (180+Δ) degrees, where Δ is an angular width of the fan beam provided by the X-ray source, to acquire a projection of a slice of the region. In some CT-imagers a “spiral scan” of a patient is performed in which the region of the patient is steadily advanced through the gantry while the X-ray source simultaneously rotates around the patient and projections of slices in the region are acquired “on the fly”.
In the above discussion it is tacitly assumed that a CT-imager images a single slice of a patient at a time. However, a modern CT-imager is very often a multislice imager that simultaneously images a plurality of slices. Such an (multislice) imager comprises a detector array having a plurality of substantially contiguous rows of detectors and the fan beam of the CT-imager is made sufficiently “thick” to illuminate all the rows of detectors. As a result, at any given view angle the CT-imager simultaneously acquires data for a number of slices equal to the number of rows in its detector array. For simplicity of presentation it is generally assumed in the discussion below that a CT-imager is a single slice imager having a detector array comprising a single row of X-ray detectors.
Determining values for the absorption integral from signals generated by X-ray detectors of a CT-imager generally requires performing a calibration procedure in which response of the CT-imager's detectors to X-rays from the imager's X-ray source is measured when nothing is located between the detectors and the X-ray source. Such a calibration procedure is referred to as an “air-calibration”.
Let a signal generated by the n-th detector responsive to incident intensity I(n,θ) be represented by SI(n,θ)=g(n,θ)I(n,θ)=g(n,θ)Ioexp(−∫μ(l)dl) where g(n,θ) is a proportionality coefficient, hereinafter referred to as a “gain”, of the detector. Let a signal generated by the CT-imager's reference detector responsive to X-ray intensity Io provided by the X-ray source be represented by RSIo=grIo, where gr is a gain of the reference detector. (In general, gr of the reference detector is independent of view angle and for convenience this is assumed to pertain in the present discussion so that gr is written as independent of θ.) The absorption integral is determined from the log of {SI(n,θ)/RSIo}. In particular, ln {SI(n,θ)/RSIo}=ln {[g(n,θ)Ioexp(−∫μ(l)dl)]/[grIo]}=[ln {g(n,θ)/gr}−∫μ(l)dl], so that the absorption integral ∫μ(l)dl=[ln {SI(n,θ)/SIo}−ln {g(n,θ)/gr}].
From the expression for the absorption integral it is seen that to determine a value of the absorption integral from the signals SI(n,θ) and RSIo the log of the gain ratio g(n,θ)/gr must be determined. Values for the gain ratio g(n,θ)/gr of detectors in the detector array are provided by performing an air calibration. During an air calibration, since there is nothing between the X-ray source and the detector array so that ∫μ(l)dl≅0 and ln {SI(n,θ)/RSIo}=[ln {g(n,θ)/gr}−∫μ(l)dl]=ln {g(n,θ)/gr}.
Generally, in an air-calibration, calibration data (i.e. signals SI(n,θ) and RSIo with nothing but air between the X-ray source and detector array) is acquired for a plurality of different view angles θ. The acquired data is processed to provide for each detector an average over a plurality of view angles of the log of the air-calibration gain ratios. It is convenient to represent the view angle average of the logarithm of the air-calibration gain ratio, hereinafter a “gain ratio factor”, for the n-th detector by the symbol AC(n) so that
      A    ⁢                  ⁢          C      ⁡              (        n        )              =                    ln        ⁢                  {                                    g              ⁡                              (                                  n                  ,                  θ                                )                                      /                          g              r                                }                    _        θ  where the overhead bar followed by θ represents an average with respect to view angle θ of the expression under the bar. For each detector, its gain ratio factor AC(n) is used to determine a value for an absorption integral from a signal SI(n,θ) provided by the detector so that ∫μ(l)dl=[ln {SI(n,θ)/SIo}−AC(n)}].
The AC(n) of a CT-imager can be dependent on values of software and hardware parameters that determine an operating configuration of the CT-imager. As a result, often values AC(n) for a CT-imager are acquired in an air-calibration of the imager for different frequently used operating configurations of the imager. For example, the AC(n) can be functions, inter alia, of an operating voltage of the X-ray source of the CT-imager and configuration of a collimator that collimates X-rays from the X-ray source to determine thickness of slices that the imager images. A full air-calibration of the CT-imager may therefore provide values for gain ratio factors for different X-ray voltages and slice thickness settings.
An air-calibration of a CT-imager may require a period of time that lasts from a few minutes to more than a half-hour to be performed and is often performed only once a day before beginning an “imaging workday” in which the CT-imager is used to image patients. Whereas an air-calibration may be performed as frequently as a few times a day, performing an air-calibration once CT-imaging of patients has begun, increases imager down time and decreases patient throughput.
However, during a day's operation of a CT-imager, gain ratios of detectors in the CT-imager often change as a result of changes in sensitivity of the detectors due, for example, to temperature changes and radiation damage and changes in communication links that transmit data from the detectors mounted on the gantry rotor to the gantry's stator. Such changes can degrade images provided by the CT-imager by, for example, generating ring artifacts in the images.
To provide some measure of adjustment for changes that may occur in gain ratio factors during CT-imaging operation, often an average
      A    ⁢                  ⁢    C    =            (              1        /        N            )        ⁢                  ∑        1        N            ⁢              A        ⁢                                  ⁢                  C          ⁡                      (            n            )                              of the gain ratio factors AC(n) is determined, where N is a total number of detectors in the detector array. Each gain ratio factor AC(n) is then usually expressed as a sum AC(n)=AC+ΔAC(n) where ΔAC(n), hereinafter a “differential gain ratio factor”, is a deviation of the gain ratio factor from the average. During CT-imaging of a region of the body of a patient with a CT-imager, generally, for each view acquired of the region, some of the detectors in the imager's detector array are not shadowed by the patient's body. X-rays from the X-ray source are incident on these non-shadowed detectors without passing through the patient's body. Signals generated by the non-shadowed detectors in the view responsive to X-ray radiation incident thereon are used to determine an average value for the gain ratio factor for the non-shadowed detectors. If the average gain ratio factor for the non-shadowed detectors is statistically significantly different from average gain ratio factor AC, then the “non-shadowed” average gain ratio factor is used to update AC and thereby values of the gain ratio factors AC(n) for the view.
Whereas, values for AC(n) of a CT-imager may be corrected, as noted above, by correcting average gain ratio factors AC responsive to average gain ratios determined for non-shadowed detectors, the corrections thus made do not correct for changes in the differential gain ratio factors ΔAC(n) of the imager. As a result, quality of CT-images acquired with the imager during a day's operation can be compromised by ring artifacts and other image defects.
It would be advantageous to have a method for updating gain ratio factors AC(n) for a CT-imager during “run time” of the imager, i.e. when the imager is being used to image patients, by updating the imager's differential gain ratio factors without having to perform air-calibration of the imager during run time.